IEEE JOURNAL OF SELECTED TOPICS IN QUANTUM ELECTRONICS, VOL. 5, NO. 4, JULY/AUGUST 1999
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Optical Coherence Tomography (OCT): A Review
Joseph M. Schmitt
(Invited Paper)
Abstract— This paper reviews the state of the art of optical coherence tomography (OCT), an interferometric imaging
technique that provides cross-sectional views of the subsurface
microstructure of biological tissue. Following a discussion of the
basic theory of OCT, an overview of the issues involved in the
design of the main components of OCT systems is presented.
The review concludes by introducing new imaging modes being
developed to extract additional diagnostic information.
Index Terms— Imaging, interferometry, optical coherence tomography, speckle.
I. INTRODUCTION
T
HE APPLICATION of optical technology in medicine
and biology has a long and distinguished history. Since
the 18th century, the microscope has been an indispensable
tool of biologists. With the invention of the laser in the
early 1960’s, physicians gained a new surgical instrument.
The development of fiber optics led to the manufacture of
endoscopes that permit direct viewing of internal organs deep
in the body. In the modern clinical laboratory, new optical
technologies facilitate the chemical analysis of tissue samples
and the counting and sizing of blood cells. In spite of these
and other advances, few of the optical instruments used in
medicine today take advantage of the coherent properties of
light. Even most instruments that employ lasers, the ultimate
generators of coherent light, can be classified as incoherent
optical systems because the focused laser beam serves mainly
as a source of illumination or concentrated heat. Perhaps one
of the reasons why optical coherence tomography (OCT) has
attracted the attention of engineers and scientists working in
the photonics field is that it has the potential to become the
first diagnostic imaging technology in which coherent optics
features prominently.
OCT is a novel imaging technology that produces highresolution cross-sectional images of the internal microstructure
of living tissue [1]. Its first applications in medicine were
reported less than a decade ago [1]–[5], but its roots lie in early
work on white-light interferometry that led to the development
of optical coherence-domain reflectometry (OCDR), a onedimensional (1-D) optical ranging technique [6], [7]. Although
Manuscript received October 26, 1998; revised June 1, 1999. This work was
supported by the Hong Kong Research Grants Council under Grant HKUST
6002/97E.
J. M. Schmitt was with the Department of Electrical and Electronic
Engineering, The Hong Kong University of Science and Technology, Clear
Water Bay, Hong Kong. He is now with Mallinckrodt Corporation, Pleasanton,
CA 94588 USA.
Publisher Item Identifier S 1077-260X(99)07521-8.
OCDR was developed originally for finding faults in fiberoptic cables and network components, its ability to probe
the eye [8]–[10] and other biological tissues [11], [12] was
soon recognized. The superb optical sectioning ability of
OCT, which is achieved by exploiting the short temporal
coherence of a broadband light source, enables OCT scanners
to image microscopic structures in tissue at depths beyond the
reach of conventional bright-field and confocal microscopes.
Probing depths exceeding 2 cm have been demonstrated in
transparent tissues, including the eye and the frog embryo
[13], [14]. In the skin and other highly scattering tissues,
OCT can image small blood vessels and other structures
as deep as 1–2 mm beneath the surface [4]–[16], [17]. An
advantage that OCT has over high-frequency ultrasonic imaging, a competing technology that achieves greater probing
depths but with lower resolution [18], is the relative simplicity
and lower cost of the hardware on which OCT systems
are based.
These attractive features notwithstanding, a number of technical problems must be overcome before OCT can take its
place alongside the established diagnostic imaging modalities.
Hidden behind the deceptive simplicity of the basic OCT
system are complicated issues that relate to the generation and
interference of partially coherent optical fields and how such
fields propagate in biological tissue. Equally important are issues related to the design of practical interferometric scanning
and detection systems. Rapid progress has been made over the
last few years toward an understanding of these issues, which
has led to new technologies for extending the capabilities of
OCT. Keeping track of developments reported in journals and
conferences that serve several diverse disciplines is becoming
increasing difficult.
This review summarizes the technological advances in optical coherence tomography that have been made over the last
ten years and places these advances in the context of key
theoretical issues involved in imaging highly scattering tissue
with partially coherent light. An overview of the technical
issues involved in the design of the main components of an
OCT system is presented. The review concludes with an introduction to parametric imaging modes of OCT that have been
developed recently to acquire information about biological
tissue in addition to its anatomical structure. Regrettably, a
comprehensive summary of the technical issues of primary
interest to optical engineers and scientists left little space for
a discussion of the applications of OCT. Readers interested in
this important topic should refer to earlier reviews by Fujimoto
et al. [5], Izatt et al. [19], and Fercher [20].
1077–260X/99$10.00  1999 IEEE
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of the source, according to
(2)
is the speed of light,
is the center
where
is its complex temporalfrequency of the source, and
. According to the
coherence function with argument
is related to the power specWiener–Khinchin theorem,
as [21]
tral density of the source,
(3)
Fig. 1. Component blocks of an OCT system.
II. THEORETICAL BASIS
OF
OCT
Fig. 1 shows the basic components of an OCT system. At
the heart of the system is an interferometer illuminated by a
broadband light source. In this section the interferometer is
first stripped to its bare essentials for analysis. Complications
introduced by scattering in a tissue sample are introduced later
in the section.
It follows from this relationship that the shape and width of the
emission spectrum of the light source are important variables
in OCT because of their influence on the sensitivity of the
interferometer to the optical path difference. Section III-A
discusses the properties of broadband sources that are suitable
for use in OCT. Sources with broad spectra are desirable
because they produce interference patterns of short temporal
and
(and spatial) extent. The relationship between
can be seen clearly when both are represented by Gaussian
functions:
(4)
with
A. Low-Coherence Interferometry
The interferometer in an OCT scanner splits a broadband
and sample field . The
source field into a reference field
sample field focuses through the scanning optics and objective
lens to some point below the surface of the tissue. After
scattering back from the tissue, the modified sample field
mixes with
on the surface of the photodetector. Given the
assumption that the photodetector captures all of the light from
the reference and sample arms, the intensity that impinges on
the photodetector is
(1)
where and are the mean (dc) intensities returning from the
reference and sample arms of the interferometer. The second
term in this equation, which depends on the optical time delay
set by the position of the reference mirror, represents the
amplitude of the interference fringes that carry information
about the tissue structure. The nature of the interference
fringes—or whether any fringes form at all—depends on the
degree to which the temporal and spatial characteristics of
and
match. Thus the interferometer functions as a
cross correlator and the amplitude of the interference signal
generated after integration on the surface of the detector
provides a measure of the cross-correlation amplitude. To
facilitate the separation of the cross-correlation signal from
the dc component of the intensity, various techniques have
A few of these techniques are
been devised to modulate
discussed in Section III-C.
Under the assumption that the tissue behaves as an ideal
mirror that leaves the sample beam unaltered, the correlation
amplitude depends on the temporal-coherence characteristics
(5)
and
(6)
represents
In these equations, the half-power bandwidth
the spectral width of the source in the optical frequency
domain. The corresponding measure of the correlation width,
derived from (6), is the (free-space) correlation length, given
by
(7)
(8)
is the full-width of the coherence function at halfwhere
maximum measured in wavelength units. Other definitions
of the coherence length yield similar expressions, but with
a different constant prefactors. For example, defined as the
speed of light in the medium times the area under the squared
amplitude of the normalized temporal coherence function,
[21]. In the OCT
literature, (8) is the most common definition.
B. Tissue Scattering Models
In the derivation of the preceding expressions, the tissue was
treated as an ideal mirror or coherent reflector. Propagation
into the tissue and back was assumed to introduce a time
delay, but otherwise leave the amplitude and coherence of
the sample beam unchanged. Generally speaking, this is an
SCHMITT: OPTICAL COHERENCE TOMOGRAPHY (OCT): A REVIEW
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unrealistic assumption. However, models based on a modified
version of this assumption provide a satisfactory description
of OCT imaging in the eye and other transparent tissues
composed of weakly reflecting layers that are flat over the
cross section of the sample beam [22]–[25]. In this special
case, the interference signal can be expressed as a convolution,
(9)
represents the fraction of power reflected from
where
in the tissue; the time delay of
a layer located at position
reflected sample beam relative to the reference beam is defined
, with the position of the reference mirror
as
given by . This simple convolution model has been used in
several studies [9], [12], [26], [27] to measure the locations and
reflectivities of tissue layers in a manner analogous to pulseinterpreted as a reflectivity
echo ultrasound, with
profile analogous to an A-mode ultrasound scan. It follows
determines the axial
from (9) that the spatial extent of
width of the point-spread function; therefore, in this onedimensional model the coherence length is an appropriate
measure of the axial resolution of the OCT system.
In the more general situation in which a two-dimensional
cross-sectional image is built up from multiple axial scans,
the optical transfer function of the optics that scan the beam
and focus it into the tissue must be considered. Kempe and
Rudolph [28] give expressions for the coherent point-spread
function of interference microscopes in the single-scattering
regime that are applicable to OCT imaging systems. They
point out the equivalence between the imaging characteristics
of confocal and interferometric systems in the transverse
dimension and show that the temporal coherence of the light
source does not influence the transverse point-spread function
of microscope significantly for coherence times greater than
about 20 fs. That is, except when the bandwidth of the
source is very broad, the axial and transverse responses of
the OCT system can be treated separately. Their studies also
emphasize the impact of aberrations caused by focusing deep
into a tissue through the air-tissue interface. Methods for
correcting these aberrations, which become more severe as
the numerical aperture (NA) of the objective lens increases,
have been proposed for microscopes with a fixed focus [29],
[30], but no practical method has yet been found for correcting
aberrations dynamically in OCT systems that require focus to
be maintained over a range of depths.
Although treating biological tissue as a collection of flat
specular reflectors is convenient for theoretical analysis, the
fact remains that most tissues are optically dense and do not
conform well to this model. Soft tissue consists of a gelatinous
matrix of collagen and elastin fibers packed with cells, blood
vessels, nerves, and numerous other structures. The dimensions
of the constituents of tissue range from less than 100 nm
to more than several millimeters. When light is focused into
tissue, the inhomogeneities in the refractive index cause the
light to scatter at various angles [3], [31]. Fig. 2 categorizes
the main types of scattering interactions. In the context of
the present analysis, a fundamental questions arises: How do
these scattering interactions affect the spatial and temporal
coherence of returning sample beam? Several research groups
Fig. 2. Scattering interactions in tissue classified into four types (see inset).
The reference field Er (; t) interferes with the distorted sample field Es0 (; t)
on the surface of the detector. The temporal cross-correlation product of the
fields is integrated over the spatial domain d2 defined by the area and
collection angle of the detector. (Adapted from [3]).
have been working toward an answer to this question, which
is central to the development of a comprehensive theory of
optical coherence tomography [3], [32]–[36].
The basic aspects of attenuation of a focused beam in tissue
composed of particulate scatterers are described by the singlebackscattering model [37], which has been adapted to the
analysis of OCT [32], [38]. This model, however, accounts for
only the subset of possible scattering interactions that cause
either total loss of coherence or no loss at all (those labeled
“1” and “2” in Fig. 2). An article by Pan et al. [25] presents a
linear systems model augmented by Monte Carlo simulations
that takes a step closer to reality. It describes the loss of the
coherence of the sample beam in terms of a convolution of the
temporal coherence function of the OCT light source and the
time-resolved reflectance function of the tissue. This article
and another by Hellmuth [24] introduce the concept that OCT
imaging systems respond to the discontinuities of the refractive
index structure on the scale of a single wavelength. In effect,
the coherence gate of an OCT system behaves as an optical
, with a width set by the
bandpass filter centered on at
coherence length of the source.
None of these models, however, accounts explicitly for
the two-dimensional (2-D) spatiotemporal distribution of the
backscattered field. Consequently, they give little insight into
the degradation of imaging quality that results from the
multiple scattering effects. Although the importance of multiple scattering has been confirmed by several experimental
studies carried out on tissue phantoms [33], [39], [40], an
effective measure that quantifies loss of contrast and resolution in images of real tissue has not yet been found.
To aid quantification of these effects, Lindmo et al. [41]
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developed a Monte Carlo simulation method that permits direct
visualization of the loss of focus that results from multiple
scattering. An analytical model developed by Schmitt and
Knüttel [36] formulates the mutual-coherence function of the
field backscattered from heterogeous tissue as a sum of the
single- and multiple-forward-scattering contributions. Among
the main findings that were obtained by applying this model
to a fractal description of the size distribution of scatterers
in tissue are: 1) low-angle multiple forward scattered light
that passes through the temporal coherence gate degrades
resolution and contrast, whereas light scattered at wide angles
degrades contrast without affecting resolution significantly;
in tissue is limited by
2) the maximum probing depth
the single-scattering coefficient and the mean scattering angle
of the tissue. Estimates of these variables for skin and other
–1.5 mm (5–8 meanhighly scattering tissues yield
free scattering lengths) at a wavelength of 1300 nm, a range
that overlaps values estimated in other studies [28]; and 3)
The coherence time of the source has a strong influence on the
resolution of an OCT scanner because it determines the width
of the axial point-spread function as well as the fineness of
the speckle generated by multiple scattering.
(a)
C. Coherent Noise (Speckle)
Although many investigators have observed the effects of
speckle on OCT imaging, its origins are not yet understood and
only a few studies concerned with speckle reduction in OCT
have been reported [42]–[44], [46]. Fig. 3(a) is an example of
an OCT image of the skin in which the image-degrading effects
of speckle are apparent. Most models of OCT wash away
speckle by averaging the spatial properties of the tissue in the
computation of the interference signal. Such averaging is not
possible in practice because static tissue produces a stationary
speckle pattern. There is a close connection between speckle
generation and the optical bandpass response of OCT imaging
systems, which makes the signal-carrying and signal-degrading
roles of speckle hard to distinguish [46].
Ways of reducing the deleterious effects of speckle on OCT
imaging include polarization diversity, frequency diversity,
spatial diversity, and image postprocessing [47]. Most of
these techniques trade resolution for reduction of speckle
contrast. Effective speckle reduction requires widening the
optical-frequency and spatial-frequency passbands of the OCT
imaging system, which are limited by the source bandwidth
and NA of the objective, respectively.
The OCT image in Fig. 3(b) illustrates the extent of speckle
reduction that can be achieved by applying the spatial-diversity
technique, while at the same time widening the source bandwidth and NA to compensate resolution losses. Features in
Fig. 3(a) are much easier to resolve than those in Fig. 3(b),
because the speckle in this image has a finer grain size and
lower contrast.
Relatively little effort has been devoted to digital processing
of OCT images for speckle reduction. Xiang et al. applied
wavelet filters that incorporate automatic noise thresholds in
an attempt to reduce high-frequency speckle noise without
blurring edges [43]. Yung et al. showed that envelope distor-
(b)
Fig. 3. Example of the effectiveness of speckle noise reduction achieved by
spatial diversity processing in combination with source bandwidth broadening.
(a) Image of skin formed from the incoherent sum of the envelopes of
0.2 detectors and two sources
interference signals produced by four NA
(source bandwidth 80 nm for 0
1280 nm). (b) Image of the same
region of skin formed from the envelope of the interference signal produced
by a single NA 0.4 detector and a single source (source bandwidth 45 nm for 0
1300 nm). Details of the apparatus and procedures used to
obtain these images can be found in [52].
=
=
1 =
=
=
1 =
tion caused by coherent interference noise can be corrected
by processing interference signals in the complex domain
[44]. However, this technique appears to work effectively only
when the number of scatterers in the sample volume is small.
Although not designed specifically to reduce speckle noise,
deconvolution methods developed by Kulkarni et al. [27] and
Schmitt [45] reduce the effects of sidelobe interference which
also contributes to coherent noise in OCT images.
III. HARDWARE
This section presents an overview of the issues involved
in the design of the optical components of the OCT system
(Fig. 1). Since most of the technology employed in OCT is
borrowed from the optical communications, it is not surprising
SCHMITT: OPTICAL COHERENCE TOMOGRAPHY (OCT): A REVIEW
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that developments in this field have quickly become incorporated into OCT systems. The design of the hardware of OCT
scanners is far from mature; therefore, this section provides
only a snapshot of progress in this area.
TABLE I
SHORT-COHERENCE SOURCES SUITABLE FOR USE IN OCT SYSTEMS
A. Light Source
The results of several theoretical and experimental studies
cited in Section II underscore the influence of the light source
on the performance of the OCT system. The general requirements of sources for OCT imaging are: 1) emission in the
near infrared; 2) short temporal coherence length; and 3) high
irradiance.
The first requirement, emission in the near infrared, stems
from the need to operate in a spectral range in which the
penetration of light into tissue is adequate. Because of the short
mean scattering length of photons in tissue at wavelengths in
the blue and ultraviolet, OCT imaging with a source that emits
in these spectral regions would be limited to superficial layers
less than a few hundred micrometers thick. At wavelengths
greater than 2500 nm, vibrational absorption by water limits
imaging to similar depths. In studies carried out to date, the
deepest penetration has been achieved using sources that emit
at wavelengths between 1200 and 1800 nm [5], [48]. The
optimum source wavelength for a given application may not
be determined entirely by penetration depth. In particular,
backscatter contrast and optical absorption (see Section IVC) are wavelength-dependent variables that may also play a
role in determining the contrast of OCT images.
The basis for the second requirement, short coherence
length, is the relationship between the temporal coherence
function of the light source and the width of the axial pointspread function of an OCT scanner (9). In general, the broader
the emission bandwidth of the source, the better the resolution
and contrast that can be achieved. However, to achieve the
benefits of broadband operation of an OCT scanner, care must
be taken to compensate for mismatches between the optical
dispersion of the reference and sample beams [49] and the
chromatic aberration of the focused beam caused by scattering
in the tissue [39]. Also, the shape of the spectrum of the source
is an important variable because it affects the dynamic range
of the scanner near strong reflections [50].
The third requirement, high irradiance, is based on the
need for wide dynamic range and high detection sensitivity
for imaging weakly backscattering structures deep in tissue.
Fortunately, because the interference signal is proportional to
the square root of the power reflected from the target (9),
a dynamic range exceeding 90 dB is not difficult to attain
with a source power greater than a few hundred microwatts.
The signal-to-noise ratio (SNR) of an interferometric imaging
system is proportional to the source power concentrated in
a single mode, which is limited by the requirement that the
of the source and its half-angle of
product of the diameter
emission be of the order of a wavelength or less [51]
(10)
This requirement is even more stringent for wideband sources.
Therefore, the most effective sources for OCT imaging emit
light from a small spot over a wide angle or from a large spot
over a narrow angle.
Table I lists characteristics of a variety of light sources suitable for use in OCT systems. The most commonly used sources
are edge-emitting light-emitting diodes (ELED’s) and superluminescent diodes (SLD’s) with peak emission wavelengths
in either the 800- or 1300-nm fiber-optic telecommunications
bands. Because of their high irradiance and relatively low
cost, superluminescent sources come close to being the ideal
sources for OCT imaging. However, the coherence lengths
of SLD’s (typically 15–30 m) are not short enough to
achieve the resolution required for many medical applications.
ELED’s with coherence lengths half as long are available
commercially at low cost, but their emission powers are
smaller by an order of magnitude. One way to overcome
the tradeoff between the source power and bandwidth is
to synthesize a broadband source by combining the outputs
of several SLD’s with different center wavelengths [52].
Multiple quantum-well devices achieve this synthesis by coupling the output of several sources on a single substrate
[53], [54].
The source power required to maintain a given SNR increases with scanning speed. To meet the demands of the
latest generation of OCT systems with scan rates that approach the television video rate, mode-locked Ti : Al O and
Cr : forsterite lasers have been employed [55], [56]. The
high power and wide bandwidth of these lasers (Table I)
make them attractive sources for fast, high-resolution OCT
imaging. However, their lack of portability limits their use to
specialized applications. The diode-pumped superfluorescent
fiber source is a lower cost and more compact alternative that
can achieve comparable powers at bandwidths up to 80 nm.
Table I gives the characteristics of three types of doped fibers
that emit at wavelengths between 1000 and 1900 nm. Wider
band operation can be achieved by splicing fibers doped with
different materials [57].
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(a)
(b)
(c)
(d)
(e)
(f)
Fig. 4. Examples of interferometer configurations suitable for use in OCT systems. (a) Standard fiber-optic Michelson interferometer. (b) Balanced
3 fiberoptic coupler [58]. (c) Balanced interferometer based on a pair of 2
2 fiberoptic couplers (An interferometer with
interferometer based on a 3
a configuration similar to this was used in 98). (d) Free-space equivalent to the fiber-optic Michelson interferometer. (e) Compact in-line interferometer
[59]. (f) Free-space Mach–Zehnder interferometer.
2
B. Interferometer
The most common interferometer configuration employed
today in OCT systems is the fiber-optic Michelson interferometer illustrated in Fig. 4(a). In this type of interferometer, light
from the source is conducted through a single-mode fiber to
an evanescent-mode coupler where half of the optical power
is extracted by another single-mode fiber that conducts the
light to the reference mirror. The remaining half of the light
enters the sample. The distal end of the fiber in the sample
arm serves a dual role as a coherent light transceiver and
spatial filter analogous to a confocal pinhole. A disadvantage
of this configuration is that the dc signal and intensity noise
generated by the light from the reference arm add to the
interference signal. In balanced configurations, examples of
which are shown in Figs. 4(b) and (c), these background noise
components are cancelled by subtracting the photocurrents
generated by two photodetectors (this subtraction can be done
by placing the terminals of two photodiodes in opposition).
The interference signals add at the output of the detectors
because they vary out of phase [58]. The Mach–Zehnder
configuration in Fig. 4(c) can be used either in the reflection
or transmission modes.
The use of fiber optics is convenient, but not essential,
in the design of OCT systems. In fact, in many applications free-space interferometers have distinct advantages. The
availability of a wide variety of prisms and mirrors give the
designer flexibility in layout of the interferometer, especially
with respect to how light is coupled into and out of the
reference and sample arms. Fig. 4(d)–(f) show a few of the
many possible configurations of free-space interferometers
suitable for use in OCT systems. The configuration in Fig. 4(d)
is the equivalent of the fiber-optic Michelson interferometer,
with the 2:2 fiber coupler replaced with a beamsplitter cube;
2
Fig. 4(f) is the equivalent of the fiber-optic Mach–Zehnder
configuration. As Fig. 4(e) illustrates, free-space optics can
be made compact for use with semiconductor sources and
detectors in handheld scanners [59]. Reducing polarization and
dispersion mismatches between the reference and sample arms
is also made easier in free-space interferometers. Because most
single-mode optical fibers are made of silica, a birefringent
material, changes in the polarization of the beam can be
induced by bending; in addition, OH absorption in the silica
limits the dispersion-free spectral width. These complications
can be avoided in free-space interferometers.
C. Beam Scanning Optics
A difficult technical problem in the design of OCT systems
is how to scan optical pathlength in the reference arm of the
interferometer rapidly and precisely. The pathlength must be
varied over a distance large enough to cover the desired axial
imaging range, which may be as large as a centimeter or more
for ocular imaging and 2 mm for imaging skin and other
optically dense tissue, and its positioning inaccuracy must be
a fraction of the source coherence length.
Most OCT systems reported to date rely on mechanical
scanning mechanisms. Fig. 5 illustrates the basic layout of
a few popular scanning methods. The simplest is based on
translation of a reference mirror mounted on a stage driven by
a dc motor or voice coil. A constant velocity is maintained
in the middle of the scan range by feedback control. The
constant-velocity movement of the reference mirror shifts the
center frequency of the interference signal to the Doppler
which facilitates removal of the dc
frequency
background and low-frequency noise during demodulation.
The maximum sustained scanning speed that can be achieved
with this approach is about 40 mm/s at a repetition rate of 30
SCHMITT: OPTICAL COHERENCE TOMOGRAPHY (OCT): A REVIEW
1211
(a)
(b)
(c)
(d)
(e)
(f)
Fig. 5. Examples of techniques for scanning the reference pathlength in OCT systems. (a) Linear translating mirror. (b) Piezo-actuated optical delay-line
constructed from parallel reflecting plates [30]. (c) FT optical delay line based on a lens-grating combination and an oscillating mirror [62]. (d) Free-space
interferometer that uses a retroreflector in the reference arm and a moving objective to achieve dynamic focusing [52]. (e) Scanner that requires only
movement of the tip of the sample fiber to achieve lateral and axial scanning, with the position of the reference mirror fixed [66], [67]. (f) Scanner
based on the optical layout of a Mirau correlation microscope [68].
Hz. Comparable scanning speeds can be attained by using a
piezoelectric transducer instead of a motor to drive a parallel
mirror system in which light reflects multiple times [Fig. 5(b)]
[17]. The development of faster mechanical scanning methods
has been the subject of numerous recent studies [60]–[65].
To date, the fastest scan speeds and repetition rates (about
175 m/s at 25000 scans/s) have been achieved by an optical
system based on a rotating beamsplitting cube driven by an air
turbine [63]. The system illustrated in Fig. 5(c) can achieve
scan rates almost as fast [64], [65] using readily available
components. Designed originally for shaping femtosecond
pulses, this system employs a grating-lens combination and
an oscillating mirror to form an optical delay line [62].
All of the techniques discussed thus far provide a way to
scan the reference pathlength, but fail to adjust the position
of the objective lens in the sample arm to keep the focus
on the optical path matchpoint. As a consequence, the axial
scan range is limited to the Rayleigh range of the focused
beam. Most researchers get around this problem by limiting
the numerical aperture NA of the scanning optics to obtain
a Rayleigh range of approximately 1 mm. At
nm, the corresponding focal spot diameter is about 30 m
for a Rayleigh range of this length. This spot size is too
large for many high-resolution imaging applications. Several
investigators have developed techniques that overcome this
problem by scanning the reference pathlength and the position
of the focused sample beam together [52], [59], [61], [66].
Fig. 5(d)–(f) show the optical layouts on which three such
designs are based. Placing the reference mirror and objective
lens on the same stage, as in the layout shown in Fig. 5(d),
permits the use of a lens with an NA as large as 0.5 [52].
This layout provides for lateral and axial scanning of the
beam and corrects the primary defocusing aberration caused
by the refractive-index mismatch at the surface of the tissue.
Designed for fiber-based OCT systems, the layout in Fig. 5(e)
requires movement of the tip of the sample fiber only [66],
[67]. It also provides for lateral and axial beam scanning and
corrects the primary defocusing aberration. The layout shown
in Fig. 5(f) is actually a self-contained interferometer based
on the Mirau arrangement that can accommodate objectives
with NA’s as large as 0.8 [68]. Its drawbacks are that it
does not correct defocus aberrations and special fabrication
methods are required to make the miniature beamsplitter and
mirror.
In an ideal OCT system, scanning of the reference pathlength and sample beam would require no moving parts at all.
In principle, coherence tomography can be carried out by electronically recording the two-dimensional interference pattern
formed at multiple optical delay times. This approach, which
falls in the domain of 3-D electronic holography [69]–[71], has
become an active area of OCT research [72]–[76]. An early
report described a depth-profiling low-coherence reflectometer
based on a stationary Fourier-transform spectrometer to which
an acousto-optical scanner was later added to allow scanning
in both the transverse and depth dimensions [72]. Fercher et
al. [74] detailed the principles and application of spectral
interferometry to the measurement of intraocular distances.
In this technique, the single photodetector of the Michelson
interferometer is replaced with a diode-array spectrometer
that records a correlogram with the reference mirror fixed
in position. Fourier transformation of the correlogram yields
the reflectivity profile of the tissue in the depth dimension.
More recently, Häusler and Lindner [76] developed a form
of spectral interferometry called “spectral radar” that can be
implemented simply by coupling a diode-array spectrometer
to the output of a fiber-optic interferometer. They used such a
system to obtain cross-sectional OCT images of the skin with
a handheld scanner.
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IEEE JOURNAL OF SELECTED TOPICS IN QUANTUM ELECTRONICS, VOL. 5, NO. 4, JULY/AUGUST 1999
Present limitations of array detectors impose the main
barriers to the development of no-moving-parts OCT scanners.
Currently available 2-D arrays that operate at near-infrared
wavelengths greater than 1000 nm are expensive and have
relatively low dynamic ranges and readout rates. The use of a
1-D photodiode array with parallel readout for one scanning
axis and a mechanical scanning apparatus for the other axis
may provide a compromise solution.
IV. NEW IMAGING MODES FOR CONTRAST ENHANCEMENT
As a noninvasive imaging method, optical coherence tomography must rely on the intrinsic variation of tissue properties to
differentiate tissue constituents. In the majority of applications
of OCT, the spatial variation of the coherent backscatter cross
section is the primary source of contrast. In principle, however,
any physical property that alters the amplitude, phase, or
polarization of the sample beam can be used to extract
information of diagnostic value. In this context, recent efforts
have turned to exploring ways of exploiting new contrast
mechanisms. In the last few years, four new OCT imaging
modes have been demonstrated: 1) polarization; 2) Doppler;
3) absorption; and 4) elasticity. A brief review of each mode
is given in the following paragraphs under separate headings.
A. Polarization
The electric fields that interfere to form the correlation
signal measured in optical coherence tomography are actually
vector quantities. For polarized sample and reference beams,
the term in (1) that represents the fringe amplitude is the
cross-correlation of the vector fields
and
It follows that the interference
signal measured by the low-coherence interferometer of an
OCT system contains time (depth)-resolved information about
the relative polarization states of the reference and sample
beams.
Muscle, tendons, and other body tissues contain collagen
and elastin fibers that exhibit birefringence when aligned
in layers. Early in the development of OCT, polarization
optics were added to the basic OCT system to permit depthprofiling of tissue birefringence [77]. The measurement of
one-dimensional birefringence profiles was later extended to
birefringence imaging of normal and thermally damaged soft
tissue [78]–[80]. Further investigations using polarizationsensitive OCT revealed that the light backscattered from
structures deep inside living skin is highly depolarized [81].
The depolarization was attributed to multiple scattering and to
single-scattering from nonspherical particles [81].
The dependence of the Stokes parameters of a random
media on the density, sizes, and arrangement of scatterers
suggests that the full characterization of the polarization state
of backscattered light may reveal information about tissue
structure that is not apparent in OCT images obtained using
randomly polarized light [82], [83]. Although the ability of
polarization-sensitive OCT to enhance the contrast between
healthy and pathological tissue has not yet been established
in clinical trials, local changes in the the degree and state
of polarization of the sample beam have been observed in
laboratory studies of thermally damaged skin [80].
B. Doppler
For nearly two decades, laser Doppler velocimetry has been
applied in a variety of medical studies, which include the
measurement of blood flow in the skin, eye, and other organs
[84], [85]. Because the laser has a long coherence length, interference between the static and Doppler-shifted components
of the light scattered in tissue occurs over an extended optical
path. Therefore, only coarse adjustment of the size and location
of the sample volume probed by the velocimeter is possible
via selection of the emission wavelength of the laser and the
geometry of the probes.
The idea of employing coherence gating to define the
sampling volume of a optical velocimeter was first applied
to the measurement of flowing particles for the study of fluid
mechanics [86]. Chen et al. took advantage of the coherencegating ability of OCT to obtain images of the velocity profile
of polystyrene beads in a tissue phantom [87] and blood flow
in a rat skin model [88]. Izatt et al. pushed the technology
of Doppler OCT further by developing methods to improve
accuracy, speed, and sensitivity [89]–[91]. Boas [92] reported
a novel method for measurement of the spatial variations in the
Brownian motion of dense concentrations of scatterers with
Doppler OCT.
Doppler OCT imaging has been demonstrated on tissue
phantoms, embryos, and small animal skin-flap models [87],
[88], [90]. Several technical challenges must be met before
Doppler OCT can achieve sufficient accuracy over the range
of depths required for imaging blood flow in human skin and
other organs noninvasively. At present, the maximum probing
depth of the OCT system limits imaging to blood vessels close
to the surface of organs. Speckle noise modulates the spectrum
of the backscattered signal, thereby reducing the signal-tonoise ratio of the centroid calculation from which estimates
of the Doppler shift are derived [91]. Moreover, a tradeoff
exists between the precision of velocity estimates and the
image frame rate [91].
C. Absorption (Spectral)
The ability of OCT to measure the absorption spectrum of a
sample can be seen by expressing the cross-correlation signal
in (1) as a product in the frequency
domain [24], [94]
(11)
(12)
(13)
is the spectrum of the source and
and
where
are the amplitude and phase spectrum of the sample, respectively. Assuming that only single-scattering and absorption
occur in the sample and that the real part of the refractive
index changes negligibly over the range of optical depths 0
the amplitude spectrum and the absorption and
to
scattering spectra at a given depth in the sample,
SCHMITT: OPTICAL COHERENCE TOMOGRAPHY (OCT): A REVIEW
and
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are related in the following way:
(14)
is the backscattering coefficient of the sample.
where
For a uniformly absorbing and scattering sample in which
scattering losses are negligible, this equation reduces to
(15)
from which an expression for the absorption spectrum is
obtained,
(16)
(17)
Although the assumptions that lead to this result are unrealistic
for highly scattering tissue samples at the wavelength of
operation of current OCT systems, (17) embodies the basic
mathematical operations that underlie spectral imaging with
OCT. Provided that the influence of the scattering coefficients
and
can be eliminated, the absorption spectrum is
proportional to the ratio of the logarithms of the normalized
Fourier-transformed cross-correlation signals measured at two
different depths in the medium. In principle, even the frequency dependence of the phase of the amplitude spectrum
that results from optical dispersion in the sample can be
extracted from the phase spectrum of the cross-correlation
signal [94]. Quantitative measurement of local variations in
tissue dispersion, especially in the eye, may have diagnostic
value.
In a recent study, the feasibility of spectral imaging was
demonstrated in tissue phantoms using an OCT system illuminated with two LED sources, one with its peak emission
wavelength centered on a strong absorption band of water
nm) and the other with its peak emission outside
nm) [93]. To reduce the influence
this band
of scattering, the ratios of the logarithms of the squared
and
were measured. However,
correlation magnitudes at
because of the relatively low absorption losses relative to
scattering losses at these wavelengths, the accuracy of the
differential-absorption method still suffers from noise caused
by speckle and spatial variations in the scattering coefficient.
As the compound in tissue with the strongest absorption
in the near infrared, water is the natural compound to target
for enhancing contrast in OCT imaging of biological tissue.
Operation of an OCT system close to the edge of the strong OH
absorption band centered on 1920 nm where the absorption and
scattering coefficients of tissue are expected to have similar
magnitudes may allow differentiation of tissue constituents on
the basis of their fat, protein, and water contents. Spectral
imaging OCT depends heavily on the characteristics of the
source. Semiconductor diode sources that emit at wavelengths
within strong water absorption bands are not available from
commercial suppliers of fiberoptic components and must be
custom manufactured. Laser-pumped superfluorescent fibers,
such as the Tm-doped fiber (Table I), which emits close to the
edge of the water absorption band that peaks at about 1920
nm, may be a suitable alternative to semiconductor diodes for
absorption imaging.
D. Elasticity
The sensitivity of optical coherence tomography to displacement of living tissue during imaging is a troublesome
problem that can lead to undesired blurring of images. OCT
elastography takes advantage of this sensitivity to quantify
microscopic deformations inside tissue induced by externally
applied stress [95]. The goal of OCT elastography or elasticity
imaging is to measure the local variations of the stiffness inside
a tissue noninvasively. The primary quantitative measure of the
stiffness is the shear elastic modulus, which varies widely for
different types of tissue [96]. A number of disease processes,
including edema, fibrosis, and calcification, alter the elastic
modulus of the extracellular tissue matrix.
At the time of this review, only one exploratory study of
OCT elastography has been carried out [95]. In this study,
stress was applied to the tissue with a piezoelectric transducer
attached to a glass plate through which the sample beam was
focused. A sequence of A-line scans was taken in either the
frame or line-by-line modes as the applied stress was increased. To generate cross-sectional images of the deformation
and strain inside the tissue, the recorded interference signals
were processed using computational methods borrowed from
ultrasonic elastography. Using this procedure, deformation
patterns in gelatin tissue phantoms, pork meat specimens,
and intact skin were obtained with a resolution of a few
micrometers.
Possible medical applications of OCT elastography include
differentiation of hard and soft masses during tissue biopsy,
imaging of arterial plaque composition, and evaluation of
wound healing. OCT elastography may also aid efforts to
solve fundamental problems in biomechanics, such as the
role of microscopic deformations in the development of the
embryo [97].
V. CONCLUDING REMARKS
Fueled by the explosion of innovations in the photonics
field, the development of optical coherence tomography (OCT)
has progressed rapidly over a period of less than 10 years.
Whether this pace of development can be sustained and
eventually lead to widespread use of OCT as a medical
diagnostic technique depends on the ability of researchers to
solve a number of tough problems that currently limit the
performance of OCT systems. More powerful and broader
band light sources, novel interferometer configurations, faster
scanning techniques, and high-contrast imaging modes are a
few of the ongoing developments that may lead to the solution
of these problems. More effort is needed to put the theory of
optical coherence tomography of biological tissue on a firmer
foundation. As the history of other medical imaging modalities
illustrates, the combination of solid theoretical analysis and
new technology offers the best hope for success.
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Joseph M. Schmitt received the B.S. degree in
biomedical engineering from Case Western Reserve
University, Cleveland, OH, in 1977 and the M.S.
and Ph.D. degrees in electrical engineering from
Stanford University, Stanford, CA, in 1981 and
1986, respectively.
He was an Associate Professor in the Department
of Electrical Engineering and Co-Director of the
Center for Medical Diagnostic Technology at the
Hong Kong University of Science and Technology,
Clear Water Bay, Hong Kong. In January 1999, he
joined the Respiratory Division of Mallinckrodt Corporation, Pleasanton, CA.